Setting laser power for laser machining stents from polymer tubing

ABSTRACT

Laser machining polymer tubing sections to form stents such that the quality and dimensions of stents from the different tubing sections are sensitive to the laser power is disclosed. The average power of the laser machining is the same for each tubing section which yields stents with different quality and strut widths.

This is a continuation application of application Ser. No. 12/554,589filed on Sep. 4, 2009, which is incorporated by reference herein.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to laser machining tubing to form stents.

2. Description of the State of the Art

This invention relates to laser machining of devices such as stents.Laser machining refers to removal of material accomplished through laserand target material interactions. Generally speaking, these processesinclude laser drilling, laser cutting, and laser grooving, marking orscribing. Laser machining processes transport photon energy into atarget material in the form of thermal energy or photochemical energy.Material is removed by melting and blow away, or by directvaporization/ablation.

The application of ultrashort-pulse lasers for high quality lasermaterial processing is particularly useful due to the extremely highintensity, ultrashort-pulse duration (<1 picosecond), and non-contactnature of the processing. Ultrashort pulse lasers allow precise andefficient processing, especially at the microscale. Compared withlong-pulse lasers and other conventional manufacturing techniques,ultrashort pulse lasers provide precise control of material removal, canbe used with an extremely wide range of materials, produce lower thermaldamage, and provide the capability for very clean small features. Thesefeatures make ultrashort-pulse lasers a promising tool formicrofabrication, thin film formation, laser cleaning, and medical andbiological applications.

However, laser machining of a substrate tends to result in unwanted heattransfer to a substrate resulting in a heat affected zone. The heataffected zone is a region on the target material that is not removed,but is affected by heat due to the laser. The properties of material inthe zone can be adversely affected by heat from the laser. Therefore, itis generally desirable to reduce or eliminate heat input beyond removedmaterial, thus reducing or eliminating the heat affected zone.

One of the many medical applications for laser machining includesfabrication of radially expandable endoprostheses, which are adapted tobe implanted in a bodily lumen.

An “endoprosthesis” corresponds to an artificial device that is placedinside the body. A “lumen” refers to a cavity of a tubular organ such asa blood vessel.

A stent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices, which function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels.

“Stenosis” refers to a narrowing or constriction of the diameter of abodily passage or orifice. In such treatments, stents reinforce bodyvessels and prevent restenosis following angioplasty in the vascularsystem. “Restenosis” refers to the reoccurrence of stenosis in a bloodvessel or heart valve after it has been treated (as by balloonangioplasty, stenting, or valvuloplasty) with apparent success.

The treatment of a diseased site or lesion with a stent involves bothdelivery and deployment of the stent. “Delivery” refers to introducingand transporting the stent through a bodily lumen to a region, such as alesion, in a vessel that requires treatment. “Deployment” corresponds tothe expanding of the stent within the lumen at the treatment region.Delivery and deployment of a stent are accomplished by positioning thestent about one end of a catheter, inserting the end of the catheterthrough the skin into a bodily lumen, advancing the catheter in thebodily lumen to a desired treatment location, expanding the stent at thetreatment location, and removing the catheter from the lumen.

In the case of a balloon expandable stent, the stent is mounted about aballoon disposed on the catheter. Mounting the stent typically involvescompressing or crimping the stent onto the balloon. The stent is thenexpanded by inflating the balloon. The balloon may then be deflated andthe catheter withdrawn. In the case of a self-expanding stent, the stentmay be secured to the catheter via a retractable sheath or a sock. Whenthe stent is in a desired bodily location, the sheath may be withdrawnwhich allows the stent to self-expand.

The stent must be able to satisfy a number of mechanical requirements.First, the stent must be capable of withstanding the structural loads,namely radial compressive forces, imposed on the stent as it supportsthe walls of a vessel. Therefore, the stent must possess adequate radialstrength and rigidity. Radial strength, which is the ability of a stentto resist radial compressive forces, is due to strength around acircumferential direction of the stent.

Once expanded, the stent must adequately maintain its size and shapethroughout its service life despite the various forces that may come tobear on it, including the cyclic loading induced by the beating heart.For example, a radially directed force may tend to cause a stent torecoil inward. Generally, it is desirable to minimize recoil. Inaddition, the stent must possess sufficient flexibility to allow forcrimping, expansion, and cyclic loading without fracturing that wouldadversely affect stent performance. Finally, the stent must bebiocompatible so as not to trigger any adverse vascular responses.

The structure of a stent is typically composed of scaffolding thatincludes a pattern or network of interconnecting structural elementsoften referred to in the art as struts or bar arms. The scaffolding canbe formed from wires, tubes, or sheets of material rolled into acylindrical shape. The scaffolding is designed so that the stent can beradially compressed (to allow crimping) and radially expanded (to allowdeployment).

Stents have been made of many materials such as metals and polymers,including biodegradable polymeric materials. Biodegradable stents aredesirable in many treatment applications in which the presence of astent in a body may be necessary for a limited period of time until itsintended function of, for example, achieving and maintaining vascularpatency and/or drug delivery is accomplished.

Stents can be fabricated by forming patterns on tubes or sheets usinglaser machining. However, as indicated above, the use of laser machiningcan have adverse effects on the properties of a material, includingpolymers.

SUMMARY OF THE INVENTION

Embodiments of the present invention include a method of fabricating aplurality of stents, comprising: providing a plurality of polymer tubingsections that are each formed separately by the same type of processingsteps for use in forming stents of the same design, wherein each of thetubing sections are made of the same polymer and have the same wallthickness, selecting an average laser power level for use in formingstents from each of the tubing sections with laser machining; and lasermachining stent patterns into the tubing sections to form a plurality ofstents using the selected laser power level, wherein the stent patternscomprise a plurality of struts, wherein the quality and dimensions ofstents obtained from the laser machining of the different tubingsections are sensitive to the laser power such that laser machining withthe selected power level that is the same for the different tubingsections yields stents with different quality and strut widths, whereinthe sensitivity is due to variations in morphology between the differenttubing sections arising from slight differences in conditions of theprocessing steps.

Embodiments of the present invention include a method of fabricating aplurality of stents, comprising: laser machining a plurality of polymertubing sections to form stents comprising a plurality of struts, whereinan average power of the laser machining is the same for each tubingsection, wherein the tubing sections are each formed separately by thesame type of processing steps for use in forming stents of the samedesign, wherein each of the tubing sections are made of the same polymerand have the same wall thickness, wherein the quality and dimensions ofstents from the different tubing sections are sensitive to the laserpower such that laser machining the different tubing sections with thesame power yields stents with different quality and strut widths,wherein the sensitivity is due to variations in morphology between thedifferent tubing sections arising from slight differences in conditionsof the processing steps.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a stent.

FIG. 2 depicts an embodiment of a portion of a machine-controlled systemfor laser machining a tube.

FIG. 3 depicts a close-up axial view of a region where a laser beaminteracts with a tube.

FIG. 4 is a plot of the strut width versus laser power for tubing madewith two different degrees of radial expansion.

FIGS. 5A-D depict images of a stent pattern formed using a laser using70 mW of power.

FIGS. 6A-C depict images of a stent pattern formed using a laser using90 mW of power.

DETAILED DESCRIPTION OF THE INVENTION

Embodiments of the present invention relating to methods of lasermachining of polymer tubing to make a stent that includes setting thepower of laser system to obtain repeatable stent dimensions. Althoughthe methods may apply to other laser machining technique, the methodsare particularly relevant to ultrashort-pulse laser machining ofsubstrates. These embodiments are suitable for fabricating fine andintricate structures of implantable medical devices such as stents.“Ultrashort-pulse lasers” refer to lasers having pulses with widths ordurations shorter than about a picosecond (=10⁻¹²). “Pulse width” refersto the duration of an optical pulse versus time. The duration can bedefined in more than one way. Specifically, the pulse duration can bedefined as the full width at half maximum (FWHM) of the optical powerversus time.

Ultrashort-pulse lasers can include both picosecond and femtosecond(=10⁻¹⁵) lasers. The ultrashort-pulse laser is clearly distinguishablefrom conventional continuous wave and long-pulse lasers (nanosecond(10⁻⁹) laser) which have significantly longer pulses. In particular,embodiments of the present method employ femtosecond lasers that havepulses shorter than about 10⁻¹³ second. Representative examples offemtosecond lasers include, but are not limited to, a Ti:sapphire laser(735 nm-1035 nm) and an excimer-dye laser (220 nm-300 nm, 380 nm-760nm).

As indicated above, embodiments of the laser machining method describedabove may be used in the fabrication of implantable medical devices suchas stents. In general, stents can have virtually any structural patternthat is compatible with a bodily lumen in which it is implanted.Typically, a stent is composed of a pattern or network ofcircumferential rings and longitudinally extending interconnectingstructural elements of struts or bar arms. In general, the struts arearranged in patterns, which are designed to contact the lumen walls of avessel and to maintain vascular patency.

An exemplary structure of a stent is shown in FIG. 1. FIG. 1 depicts astent 10 which is made up of struts 12. Stent 10 has interconnectedcylindrical rings 14 connected by linking struts or links 16. Theembodiments disclosed herein are not limited to fabricating stents or tothe stent pattern illustrated in FIG. 1. The embodiments are easilyapplicable to other stent patterns and other devices. The variations inthe structure of patterns are virtually unlimited. The outer diameter ofa fabricated stent (prior to crimping and deployment) may be between0.2-5.0 mm. For coronary applications, a fabricated stent diameter is2.5-3.5 mm. The length of the stents may be between about 6-12 mm.

The present embodiments are particular relevant to laser machiningpolymer substrates to form stents, however, the methods may beapplicable to other materials such as metals and ceramics or compositematerials composed of combinations of polymer, metal, and ceramic.Polymers can be biostable, bioabsorbable, biodegradable, or bioerodable.Biostable refers to polymers that are not biodegradable. The termsbiodegradable, bioabsorbable, and bioerodable, as well as degraded,eroded, and absorbed, are used interchangeably and refer to polymersthat are capable of being completely eroded or absorbed when exposed tobodily fluids such as blood and can be gradually resorbed, absorbed,and/or eliminated by the body. In addition, a medicated stent may befabricated by coating the surface of the stent with an active agent ordrug, or a polymeric carrier including an active agent or drug. The drugcoating is typically applied to the stent body or scaffolding afterbeing formed by laser machining. The coating is typically much thinnerthan the struts of the scaffolding, for example, the coating can be 1-5microns in thickness while the struts can be 140-160 microns thick.

An implantable medical device, such as a stent, can be fabricated bylaser machining a construct to form the device. Material is removed fromselected regions of the construct which results in formation of thestructure of the device. In particular, a stent may be fabricated bymachining a thin-walled tubular member with a laser. Selected regions ofthe tubing may be removed by laser machining to obtain a stent with adesired pattern. Specifically, a beam can be translated or scanned overthe surface of a tubing resulting in removal of a trench or kerfextending all the way through a wall of the tubing. When a starting andending point of a kerf meet, the region surrounded by the kerf drops outor is removed. Alternatively or additionally, the tube can be translatedand rotated to allow machining of tubing.

In exemplary embodiments, a stent can be cut from a tubing using amachine-controlled laser as illustrated schematically in FIG. 2. FIG. 2depicts an embodiment of a portion of a machine-controlled system forlaser machining a tube. In FIG. 2, a tube 200 is disposed in a rotatablecollet fixture 204 of a machine-controlled apparatus 208 for positioningtubing 200 relative to a laser 212. According to machine-encodedinstructions, tube 200 is rotated and moved axially relative to laser212 which is also machine-controlled. The laser selectively removes thematerial from the tubing resulting in a pattern cut into the tube. Thetube is therefore cut into the discrete pattern of the finished stent.

FIG. 3 depicts a close-up view of a laser beam 408 interacting with atube 414. Laser beam 408 is focused by a focusing lens 338 on tube 414.Tube 414 is supported by a controlled rotary collet 337 at one end andan optional tube support pin 339 at another end. A coaxial gas jetassembly 340 directs a cold gas jet or stream 342 that exits through anozzle 344 that cools the machined surface as the beam cuts and ablatesa substrate. The gas stream also helps to remove debris from the kerfand cool the region near the beam. Gas input is shown by an arrow 354.Coaxial gas jet nozzle 344 is centered around a focused beam 352. Insome embodiments, the pressure of the supplied cooling gas is between 30and 100 psi. An exemplary flow rate of the cooling gas is between 2 and10 scfh. Exemplary cooling gases or process gases include helium, argon,nitrogen, or oxygen.

It may also be necessary to block laser beam 414 as it cuts through thetop surface of the tube to prevent the beam, along with the moltenmaterial and debris from the cut, from impinging on the inside oppositesurface of tube 414. To this end, a mandrel 360 supported by a mandrelbeam block 362 is placed inside the tube and is allowed to roll on thebottom of the tube 348 as the pattern is cut. This acts as a beam/debrisblock protecting the far wall inner diameter.

The present invention is applicable to femtosecond pulsed lasers havingpulse widths of 5-10 fs, 10-80, 80-120 fs, 120-500, or 500-1000 fs. Itis also believed the invention is applicable to lasers with pulse widthsgreater than 1000 fs (1 ps), greater than 10 ps, in particular, 10-15ps.

The repetition rate used for laser machining polymers for stents with afemtosecond laser is generally between 1 and 5 kHz. The energy per pulseof such laser machining is generally 2-1000 μJ, more narrowly, 20-30 μJ.The fluence of such laser machining is generally 1-20 J/cm², or morenarrowly, 5-15 J/cm². The average power per pulse or power of a beam canbe 10-1000 mW, or more narrowly 50-150 mW. Fixed and variable wavelengthlasers may be used for laser machining polymers. Exemplary fixedwavelength lasers may have wavelengths at 248 nm, 532 nm, or 800 nm.

The embodiments of the present invention relate to determining the powerof the femtosecond laser used for laser machining a polymer tubing. Apower is determined for a tubing section that reduces or eliminatesundesirable cutting affects. The power determination also providesrepeatable stent dimensions, such as strut widths, as compared to othertubing sections processed separately prior to laser machining.

Embodiments of the method include providing a plurality of polymertubing sections that are each formed separately by the same type ofprocessing steps for use in forming stents of the same design. The samedesign refers to features such as the stent pattern and the dimensionsof the structural elements or struts of the stent pattern includingthickness and width. Other features of a design include mechanicalproperties such as radial strength and fracture toughness. Additionalfeatures include morphology properties such as polymer orientation,crystallinity, and the size of crystallites of a semicrystallinepolymer. The processing of stent precursor constructs or tubing sectionsis performed so that such features are as close as possible for tubingsections made at different times.

For each tubing section, a laser power level is determined for use informing stents from each of the tubing sections with laser machining.The method further includes laser machining stent patterns into thetubing sections to form a plurality of stents using the laser powerlevels determined for each tubing section. The stent patterns include aplurality of struts. The laser power determined for each tubing sectionis selected to obtain repeatable strut widths in the stent patternsformed from the different tubing sections.

The inventors had used a femtosecond laser (120 fs, fluence=10±5) tomachine PLLA tubes to form stents. The machining power was set to avalue that was believed to be significantly higher than a minimum valuerequired to machine through the tubing to form struts. The inventorsused the same power for all lots of tubing. A “lot” of tubing refers toa section of tubing made or processed at a given time, for example, alot of tubing extruded and then radially expanded.

Specifically, the inventors unexpectedly found that laser machiningdifferent lots with the same power with the same laser yielded stentswith different quality and strut widths. The different lots were of thesame material, PLLA, and all had the same wall thickness. The differentlots were made at different times either with the same processingconditions or only slightly different processing conditions.

The inventors observed that the quality of a machined stent and stentdimensions obtained from laser machining different lots of tubing wasextremely sensitive the laser power. “Quality” refers to several aspectsof a cut stent pattern. One aspect is the degree of smoothness of thecut. A poor quality cut can have rough portions on a cut surface, suchas flashes which refers to un-cut material that is torn from the excesscut away sections. Another aspect of a poor quality cut is “glitter”which refers to re-deposited material as particles large enough tocreate a reflective surface. Excess melting in portions of the stentpattern are another characteristic of a poor quality cut. Anothercharacteristic of a poor quality stent pattern are islands, which areportions of tubing that are not intended to be part of the stent patternthat do not fall out of the cut pattern since a region surrounded by akerf is still attached to a strut. Islands may result when the power isinsufficient for a beam to consistently cut all the way through a tubingwall as it travels around a region to be removed.

The differences in machining the different lots of tubing are asensitivity to power likely arising from variations in morphology of thepolymer in the different lots of tubing. The difference in morphologycan arise from even the slightest differences in processing conditions.Even slight differences in process conditions will likely result in suchmorphology variations between lots of tubing. However, it was unknown tothe inventors that such minor variations in processing conditions priorto machining could have such a dramatic affect on cutting quality andstrut dimensions. Inventors have found it essential to compensate forunexpected sensitivity to these variations through adjustment of thepower for different lots of tubing.

The processing referred to above includes forming a tube from a polymerresin using an extrusion process followed by radially expanding andaxial elongating the tube. A stent is formed by laser machining a stentpattern in the expanded and elongated tube. The tube is radiallyexpanded and axially elongated to increase the radial strength andfracture toughness of the tube and a stent made therefrom.

A semicrystalline polymer, such as PLLA, includes crystalline regionsseparated or surrounded by amorphous regions. Morphology includes, butis not limited to, degree of crystallinity, molecular orientation ofpolymer chains, and crystallite size. Molecular orientation refers tothe relative orientation of polymer chains along a longitudinal orcovalent axis of the polymer chains. The orientation can refer to theorientation of crystalline lamella and to the orientation of polymerchains in the amorphous regions.

With regard to making a tube by extrusion, representative examples ofextruders include single screw extruders, intermeshing co-rotating andcounter-rotating twin-screw extruders and other multiple screwmasticating extruders. For example, tubing for a stent can be formedwith a 1″ single screw extruder. A temperature range in the extruder isat least 20° C. above the Tm of the polymer. For example, an exemplarytemperature range for PLLA extrusion is 200-225° C. Other exemplaryprocessing conditions for PLLA tubing include a residence time in theextruder of approximately 10 min, quench in room temperature water bath,a die/quench distance of ¾ in, a pull rate of 16 ft/min, a barrelpressure of 2000 psi, and a draw down ratio approximately 3:1 (ID die toID of drawn tube).

A polymeric tube is radially expanded and axially elongated byincreasing the pressure inside the tube and applying a tensile forcealong the cylindrical axis of the tube, respectively. The pressureinside of the tube is increased by conveying a fluid into the tube toincrease the internal pressure in the tube. Preferably, the tensileforce is applied at one end while holding the other end stationary. Thetube is heated to a temperature between the glass transition temperature(Tg) and the melting temperature (Tm) of the polymer to allow the radialexpansion and axial elongation of the tube.

At the start of the process, the tube is positioned in a cylindricalmember or mold. The process parameters are adjusted so that the tubeexpands against the inside surface of the mold so that the outerdiameter of the expanded tube is the inside diameter of the mold. Oneend of the tube is sealed or blocked and a gas such as air, nitrogen,oxygen, argon, etc. is conveyed in the other end of the polymer tube toincrease the pressure in the tube.

The tube is heated by a heating source such as a nozzle or nozzlesblowing a warm gas onto a portion of the tube. The nozzle(s) aretranslated along the cylindrical axis of a the tube from a proximal endto a distal end, blowing warm gas onto an axial section or portion ofthe mold as it translates which heats the axial section or portion ofthe mold and the axial section or portion of the tube within the mold.The temperature and nozzle rate are adjusted so that as the nozzletranslates, the heated portion expands. The radial expansion follows thetranslating nozzle and propagates along the cylindrical axis of thetube. As the nozzle translates, the an end of the tube is pulled at aspecified rate, which is preferably constant.

The nozzle rate and pull rate are adjusted so that expansion and axialelongation start at the same time and are completed at the same time.Additionally, the nozzle rate and pull rate are preferably constantsince the properties of a deformed polymer generally depend on the rateof deformation.

Prior to the expansion and elongation, the tube is pre-heated close tothe deformation temperature. Pre-heating can be performed by a nozzletranslated along the length of the tube without the increased pressureand the tension. After the expansion and elongation is completed, thepolymer tube is cooled or allowed to cool to below its Tg either beforeor after decreasing the pressure and/or decreasing tension. Cooling thetube helps insure that the tube maintains the proper shape, size, andlength following its formation. A section of tubing expanded in the waydescribed is considered a “lot” of tubing.

Any slight variations in the processing conditions from lot to lot canresult in differences in morphology from lot to lot. Variations canresult from adjustments to processing conditions so that the expansionoccurs in the manner desired, as described above. Such slight variationsin processing conditions unexpectedly alter the power requirements forobtaining a quality machined stent with repeatable strut dimensions.

Exemplary processing conditions for a PLLA tube deformation include adeformation temperature of 75-130° C., an expansion pressure of 110-140psi, a nozzle translation rate 0.2-1.2 mm/s, a pull rate of 0.4-4.0mm/s. The percent degree of radial expansion (% RE) of a tube is100%×(Inside Diameter of Expanded Tube/Original Inside Diameter ofTube−1). The percent degree of axial elongation (% AE) is: 100%×(Lengthof Elongated Tube/Original Length of Tube−1). Exemplary % RE are200-500% and % AE are 20-200%. The degree of crystallinity of exemplaryexpanded PLLA tubes and stents made therefrom are 20-50%, or morenarrowly, 45-50%.

In addition to selecting the power for quality and repeatability, poweris also selected to reduce the heat affected zone of a machinedsubstrate. As mentioned above, the heat affected zone is a region on thetarget material that is not removed, but is affected by thermal energyfrom the laser. In the heat affected zone, the properties of material inthe zone vary as a function of distance. Stent performance may beincreased by reducing or eliminating the variation in properties of theheat affected zone. Stent performance includes having a high radialstrength sufficient maintain patency, low recoil, and resistance tocracking upon crimping and deployment.

The inventors have found that the variation of properties in the heataffected zone depends on the power. As the power increases the depth andvariation in properties changes. This has been demonstrated bynanoindentation results. Therefore, a lower power is preferable.

In some embodiments of the invention, a selected machining power formaking a stent from a selected tubing lot includes determining an aapproximate threshold power level for the laser to cut a kerf or channelall the way through the tubing wall. The threshold power can correspondto the power level that allows the laser beam to cut through a length oftubing completely when the beam in scanned over the length, in spite ofslight variations in the thickness and properties in different portionsof the tubing. A power level that results in portions along the lengththat are cut all the way through and portions along the length which arenot when the laser is scanned along the length is below the thresholdpower. The length of such a section for determining the threshold powermay be 0.2-20 mm, or more narrowly, 0.2-0.5 mm, 0.5-1 mm, 1-5 mm, 5-10mm, or 10-20 mm.

Each lot of tubing can have a different threshold power level. Anydifference or change to the material such as thickness, percent radialexpansion, percent axial elongation, or crystallinity may alter thethreshold power. The inventors have found that different lots of PLLAtubing have different threshold power levels.

In some embodiments, the threshold power may be the stent cutting powerwhich is the power used to laser machine stents from the lot of tubing.In other embodiments, the stent cutting power level may be above thethreshold power. In such embodiments, exemplary stent cutting powerlevel can be A×threshold power, where “A” is between 100% and 120%, ormore narrowly, 100% and 110%. “A” can be between 100% and 102%, 102% and105%, 105% and 108%, 108% and 112%, or 112% and 120%. In otherembodiment, A can be greater than between 120-150%, or greater than150%.

The inventors have found that a preferable stent cutting power level is110% of the threshold power level. In some embodiments, a stent patterncut using the power is free or almost free of flash, glitter, or meltedportions.

In some embodiments, the threshold power can be determined by firstselecting an initial power level that is believed to cut a kerf orchannel all the way through the tubing wall, as described above. Thepresumption is then verified by machining a length of the tubing at theinitial power level. Then, a length of tubing is machined at a powerlevel lower than the initial power and the machining portion isinspected. The process is repeated by making discrete steps downward inpower until a power level (“sub-threshold power”) is found that resultsin portions along the machined length that are cut all the way throughand portions are not after the laser is scanned over length.

In some embodiments, the lowest power level selected that is greaterthan the sub-threshold power can be selected as the threshold power. Inother embodiments, one or more additional power levels between the twolowest power levels can be tested to obtain a more accurate value of thethreshold power.

EXAMPLES

The examples and experimental data set forth below are for illustrativepurposes only and are in no way meant to limit the invention. Thefollowing examples are given to aid in understanding the invention, butit is to be understood that the invention is not limited to theparticular materials or procedures of examples.

Example Set 1

The following set of examples describes results of determining lasermachining power for four lots of PLLA tubing. Each lot of tubing wasformed from an extrusion process from 100% PLLA resin. The dimensions ofthe tubing, the extruded dimensions, are: outside diameter (OD)=0.0066inch and inside diameter (ID)=0.0025 inch. The extruded PLLA tubes wereradially expanded according the process described above. The target % REwas 400%.

Titanium-Sapphire fixed wavelength lasers were used with a wavelength of800 nm. The pulse widths of the lasers ranged from 95-120 ps with arepetition rate of 5 kHz. The fluence was between 10±5 kJ/cm². For agiven laser, it was found that each lot had a different selectedmachining power was determined. Additionally, for a given lot of tubing,the selected machining power was different for each laser. The selectedmachining powers (110%×threshold power) determined for various lots oftubing were between 90-140 mW.

Table 1 provides the selected machining powers that were determined forthe four lots of tubing for three different lasers. The deformationparameters for each lot of tubing is given in Table 2.

TABLE 1 Selected machining power for tubing determined for four lots andthree lasers. Lot 1 Power Lot 2 Power Lot 3 Power Lot 4 Power (mW) (mW)(mW) (mW) Laser 1 140 135 128 140 Laser 2 105 110 90 — Laser 3 — 130 128100

TABLE 2 Deformation parameters for lots of tubing listed in Table 1.Deformation Parameters Lot 1 Lot 2 Lot 3 Lot 4 Expand Heat (F.) 166 185178 180 Expand Air Flow (scfh) 60 60 60 60 Pre-heat Dwell (sec) 32 30 3832 Heat Nozzle Speed (mm/s) 0.32 0.30 .38 .32 Expand Pressure (psi) 110130 110 115 Number of Passes 1 1 1 1 Initial Tension Position (mm) 25 2525 25 Pre-Tension (Grams) 180 180 180 180 Set Heat (F.) 32 32 32 32 SetAir Flow (scfh) 20 20 20 20 Set Heat Dwell (min) 0 0 0 0 Cool Time (sec)30 30 30 30 Tubing Start OD (thou in) 66 66 66 66 Tubing Start ID (thouin) 25 25 25 25 Tubing End OD (thou in) 136 136 136 136 Tubing End ID(thou in) 124.2 124.2 124.2 124.2

Example Set 2

The affect of power and the degree of radial expansion on strut widthwas studied. The inventors observed that a repeatable strut width wasnot obtained when the same power was used to machine different lots oftubing. The inventors also observed that for a given power, differentstrut widths were obtained from tubing with different degrees of radialexpansion. The laser power had to be adjusted each time to obtain adesired strut width.

The affect of power and the degree of radial expansion was studied bylaser machining two PLLA tubes with two different % RE, 300% and 400%,at several power levels. The strut widths of each of the stents weremeasured to determine the affect of power on strut width. The expandedinside and outside diameter of the expanded tubes are approximately thesame. [Do you have tubing thickness for each group?]

A summary of the strut width data is shown in Table 3 and the data isplotted in FIG. 4.

TABLE 3 Laser machining data showing dependence of strut width on powerand radial expansion. % RE Power Strut width (inch) 300 80 0.00674 30090 0.00659 300 103 0.00653 300 123 0.00642 300 145 0.00636 400 700.00704 400 90 0.00693 400 120 0.00683

The data in Table 3 and FIG. 4 demonstrate that:

1. As the power increases when laser machining a given lot of tubing,the strut width decreases. The trend was apparent on tubing with twodifferent radial expansions studied.2. The degree of radial expansion affected strut width.

The two groups were cut with the same laser and the same stent patternprogram. The measurements were taken with a Keyence optical microscopeat 200×, calibrated to a pin gauge. Power dependence shows that if samepower is used for all tubes, different strut widths will be obtained.

Example Set 3

Stent patterns cut from a PLLA tube at two different power levels showedthe affect of power on the quality of cutting. The degree of radialexpansion of the PLLA tubes was 400%. The expanded tubing had a 0.136 inOD and a 0.006 inch nominal wall thickness. The pulse width of the laserwas 92 fs and the height of cooling nozzle blowing helium gas was 0.35.Exhaust was 1150 ft/min. [What is this?] Stent patterns were cut at twopower levels, 70 mW and 90 mW.

The stent patterns cut with a power level of 70 mW and an helium coolingflow gas of 4 scfh had flashes, roughness, and islands. FIGS. 5A-D areexemplary images of portions of the stent patterns cut under theseconditions. FIG. 5A illustrates an island and flash. FIG. 5B showsflash. FIG. 5C shows roughness. FIG. 5D shows flash and glitter.

FIGS. 6A-C are exemplary images of portions of the stent patterns cutwith a power level of 90 mW. FIGS. 6A-B are portions cut with heliumcooling gas flow of 4 scfh and FIG. 6C are portions cut with the flow at6 scfh. As shown in FIGS. 6A-B, the flash were no longer present,however, melt and glitter were. The helium flow was raised from 4 to 5,but the melt and glitter remained (not shown). As shown in FIG. 6C,after the flow was increased to 6 scfh, the melt and glitter were nolonger present.

As used herein, “substantially the same” or “almost the same” can referto within 0.01% to 5%.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

What is claimed is:
 1. A method of fabricating a plurality of stents, comprising: providing a plurality of polymer tubing sections that are each formed separately by the same type of processing steps for use in forming stents of the same design, wherein each of the tubing sections are made of the same polymer and have the same wall thickness, selecting an average laser power level for use in forming stents from each of the tubing sections with laser machining; and laser machining stent patterns into the tubing sections to form a plurality of stents using the selected laser power level, wherein the stent patterns comprise a plurality of struts, wherein the quality and dimensions of stents obtained from the laser machining of the different tubing sections are sensitive to the laser power such that laser machining with the selected power level that is the same for the different tubing sections yields stents with different quality and strut widths, wherein the sensitivity is due to variations in morphology between the different tubing sections arising from slight differences in conditions of the processing steps.
 2. The method of claim 1, wherein each tubing section is made of the same material and has the same wall thickness and outer diameter.
 3. The method of claim 1, wherein a pulse width, repetition rate, and fluence are the same in the laser machining of each of the tubing sections.
 4. The method of claim 1, wherein the processing steps comprise tubing extrusion and radial expansion of the extruded tubes, the method further comprising extruding a plurality of tubing sections and radially expanding the extruded tubing sections to form the plurality of tubing sections.
 5. The method of claim 1, wherein the laser is a femtosecond laser with a pulse width between 95-120 fs.
 6. The method of claim 1, wherein the tubing sections are made of PLLA.
 7. The method of claim 1, wherein the stents formed using the determined laser power levels are free of defects including flash, glitter, and melted portions.
 8. The method of claim 1, wherein the laser machining parameters include a pulse width of 95-120 fs, a repetition rate is 2.5-5 kHz, a power level of 0.2-0.3 mW, and a fluence of 5-15 J/cm2.
 9. A method of fabricating a plurality of stents, comprising: laser machining a plurality of polymer tubing sections to form stents comprising a plurality of struts, wherein an average power of the laser machining is the same for each tubing section, wherein the tubing sections are each formed separately by the same type of processing steps for use in forming stents of the same design, wherein each of the tubing sections are made of the same polymer and have the same wall thickness, wherein the quality and dimensions of stents from the different tubing sections are sensitive to the laser power such that laser machining the different tubing sections with the same power yields stents with different quality and strut widths, wherein the sensitivity is due to variations in morphology between the different tubing sections arising from slight differences in conditions of the processing steps.
 10. The method of claim 9, wherein each tubing section is made of the same semicrystalline polymer material and each tubing section has the same wall thickness and outer diameter.
 11. The method of claim 9, wherein the adjusted power for each tubing section is different from the adjusted power for other tubing sections.
 12. The method of claim 9, wherein at least one processing parameter of the processing steps of each tubing section differ and the morphology of the processed tubing sections depends on the at least one parameter.
 13. The method of claim 12, wherein the processing steps comprise tubing extrusion and radial expansion of the extruded tubes, the method further comprising extruding a plurality of tubing sections and radially expanding the extruded tubing sections to form the plurality of tubing sections.
 14. The method of claim 12, wherein laser machining is performed with a laser having a pulse width of 80-120 fs.
 15. The method of claim 12, wherein laser machining is performed with a laser having a pulse width of 10-15 ps. 